Typically, in CT imaging systems, a rotatable gantry includes an x-ray tube, detector, data acquisition system (DAS), and other components that rotate about a patient that is positioned at the approximate rotational center of the gantry. X-rays emit from the x-ray tube, are attenuated by the patient, and are received at the detector. The detector typically includes a photodiode-scintillator array of pixelated elements that convert the attenuated x-rays into photons within the scintillator, and then to electrical signals within the photodiode. The electrical signals are digitized and then received within the DAS, processed, and the processed signals are transmitted via a slipring (from the rotational side to the stationary side) to a computer or data processor for image reconstruction, where an image is formed.
The gantry typically includes a pre-patient collimator that defines or shapes the x-ray beam emitted from the x-ray tube. X-rays passing through the patient can cause x-ray scatter to occur, which can cause image artifacts. Thus, x-ray detectors typically include an anti-scatter grid (ASG) for collimating x-rays received at the detector.
Imaging data may be obtained using x-rays that are generated at a single average energy, corrected from a polychromatic energy spectrum. However, some systems may obtain multi-energy images that provide additional information for generating images.
Third generation multi-slices CT scanners are built with detectors that include scintillator/photodiodes arrays. These detectors are positioned in an arc where the focal spot is the center of a corresponding circle. The material used in these detectors generally use scintillation crystal/photodiode arrays, where the scintillation crystal absorbs X rays and converts the absorbed energy into visible light. A photodiode is used to convert the light to an electric current.
In recent years the development of volumetric (VCT) or cone-beam CT (CBCT) technology has led to a rapid increase in the number of slices used in CT detectors. The detector technology used in large coverage CT enables greater and greater coverage in patient scanning by increasing area exposed. In CT detectors, the increase of the number of slices results in an increase in the width of the detector along a length of the patient, or commonly referred to as the Z-axis.
The x-ray detectors of current state of the art CT systems include a two-dimensional (2D) array of scintillating pixels, coupled to a 2D array of Si photodiodes. A typical detector can include an array of 16, 32, 64, or more. However, recently the need for cardiac imaging has gotten more interest, with the goal of imaging the heart within one rotation. In order to image the heart in one rotation, for common gantries the detector size is therefore approximately 140 mm to 160 mm (at iso-center or “ISO”) to cover the full organ in half scan (which in one example is equivalent to a detector having approximately 256 slices).
Building very large modules in a monolithic structure to cover 160 mm in z-axis coverage includes numerous challenges, such as manufacturing cost and reliability. To reduce cost, one known method includes abutting smaller modules to have more than one module extending along the Z-axis. In such a design, each module may include its own collimator attached to the surface of the scintillator, and as indicated one or more modules may thereby be built into a larger structure, extending in a Z-direction.
However, alignment of collimators with their respective scintillator can be challenging, and may include optical alignment features to position the collimator with respect to pixels of the scintillator. Optical placement can be difficult, in that alignment of the relatively deep (in a Y-direction) of the collimator may not be possible due to the depth of the scintillator, and individual pixels of the scintillator themselves may not be visible during assembly.
In addition, adhesion to the scintillator can also be a challenge, as the scintillator is typically covered with reflector material to prevent stray light from passing to the pixels, and the reflector on the scintillator may not have sufficient surface adhesion characteristics to obtain or maintain robust and reliable adherence during the life of the detector.
Thus, there is a need to improve the scintillator adhesion to the collimator, precise alignment and assembly thereof.